Positron emission tomography (PET) can detect tumor in vivo based on the living chemistry of cancer tissues. In recent years, PET has well demonstrated its broad clinical utilities in cancer diagnosis and is recognized as an important tool to study cancer functions in vivo, because of its unique ability to elicit molecular functions. The in vivo molecular imaging ability of PET has triggered considerable cancer-research interest and radiotracer development to study cancer related molecular processes such as angiogenesis, apoptosis, cell proliferation, hypoxia, gene expression and blood flow.
Despite this success, the cancer-application potential of PET in whole body is still largely untapped today because of high scanner cost, and low image resolution. As of 2009, clinical PET cameras have imaging resolution of 4.0-6.3 mm, but because of low sensitivity, the practical clinical resolution is worse (7-10 mm), which can miss smaller (early) primary lesions and metastases. In the last decade, researchers have focused on the development of low-cost ultrahigh resolution PET technologies and PET cameras.
PET camera detectors are made up of tens of thousands of scintillation crystals and thousands of photosensors. The most commonly used photosensors in clinical PET cameras are photomultiplier tubes (PMT). Although other solid-state or semiconductor photosensors are also being investigated, such sensors are currently more expensive than PMT.
Photomultiplier tubes (PMT) are used to detect scintillation signals from scintillation crystals and convert the signal into electronic signals. In certain existing embodiments, an array or block of scintillation crystal elements (pixels) are coupled to the four or more PMT's. These four PMT's detect the light from a scintillating crystal element and decode the position of the scintillating crystal within the block. Each crystal element (pixel) distributes a unique ratio of light to each of the four PMT. The unique ratio of light distribution from each pixel to the four or more PMT is the signature of each pixel for decoding that pixel. One of the typical designs is shown in FIGS. 1 and 2.
PMT cost is a major component cost in a PET detector system. For example, in a typical clinical PET, about 1,200 PMT are used (e.g. DST PET-CT available from GE® Corporation) and each PMT channel costs $250-300 (with high voltage circuit and amplifiers). Accordingly, reducing the number of PMT used can lower the production cost of PET scanners, and making the crystal elements smaller can improve the imaging resolution of PET cameras. To reduce the number of PMT used, and to improve the image resolution of PET camera, a detector design utilizing “Photomultiplier Quadrant Sharing” (PQS) was developed as shown in FIGS. 3 and 4 (and as described in U.S. Pat. Nos. 5,319,204; 5,453,623; 6,956,214 and 7,238,943). Current exemplary embodiments of the PQS design are not confined to using photomultiplier tubes, but can also use solid state photosensors. Accordingly, the term “PQS” has been updated to mean “Photosensor-Quadrant-Sharing” in this disclosure.
In the PQS design, each quadrant of a PMT is placed adjacent to one quadrant of each detector block/array. The PQS design reduces the number of PMT used by seventy-five percent because each PMT measures the light output of four arrays instead of one array, thus replacing four PMT with one PMT. By the same token, if the same size PMT and the same number of PMT are used, the detector pixel can be reduced by seventy-five percent, quadrupling the number of pixels in the camera. Since the two dimensional data acquired is used to generate a three-dimensional tomography image, the camera imaging volume or voxel size is decreased by eight times using the PQS design. Accordingly, the PQS design may be used to detect cancer lesions that are one-eighth the size of those detectable by non-PQS designs, without increasing the production costs of the current clinical systems.
Generally, there are two methods to distribute light from each crystal pixel of a block/array to the four PMT for decoding the position of a firing crystal. One method is to have the crystal block or array coupled to a light guide with different saw-cut depths (see FIGS. 1 and 3). This light guide is then coupled to four PMT, either in the conventional way (see FIG. 1) or in the PQS configuration (see FIG. 3). A second method is to distribute light from the firing crystal through adjacent crystal pixels in the block to four PMT (see e.g., FIGS. 2 and 4). There are two approaches to implement this second method. For example, embodiments provided by GE® Corporation uses crystals with different surface finishes (on the 4 cylindrical sides) to control the amount of light crossover from one crystal to the next. However, since there are only a limited effective surface finishes that can be used, the number of crystals that can be decoded by the four PMT is more limited. The smaller number of crystal pixels in a block/array leads to lower imaging resolution. The second approach, developed by the inventors of the exemplary embodiments disclosed herein, uses partial reflectors (white paint or mirror film) applied to each surface of each pixel as shown in FIG. 5.
FIG. 1 displays a conventional PET detector array using external light guides to distribute light from scintillation detector pixels to four PMT, while FIG. 2 displays a conventional PET array using internal light distribution to distribute light to four PMT. FIG. 3 illustrates a PQS array using an external light guide and the same size PMT as FIGS. 1 and 2. The FIG. 3 configuration can provide much smaller detector pixels (higher imaging resolution), or utilize a PMT that is four times larger to provide the same size detector pixel (to reduce the number of PMT used). FIG. 4 illustrates a PQS array using internal light distribution to increase resolution or reduce PMT cost.
Partial reflectors can be used for controlling a desirable amount of light bleeding to a neighboring crystal pixel. FIG. 5 illustrates the partial reflectors (e.g., paint) applied to the detector pixels in a block. Since the partial reflector can have infinite sizes and shapes, many more crystals pixels can be decoded by four PMT, leading to higher resolution. Inventors of exemplary embodiments disclose herein have also invented and patented a manufacturing method (Slab-Sandwich-Slice or SSS method) to efficiently and accurately produce position-decoding blocks. Again, this partial reflector block-detector design can be coupled to four PMT conventionally (FIG. 2) or in PQS configuration (FIG. 4).
PQS can therefore either be used to lower the cost of nuclear camera production cost or improve the imaging resolution of cameras, or both. With continuous design improvement on the internal structure of PQS detectors, the inventors have achieved much more than four times higher detector pixel resolution. Exemplary embodiments of the PQS detectors developed by the inventors can decode 256 detectors per photosensor, while the current GE® Corporation PET camera decodes 14 detectors per photosensor. The current Siemens® PET camera decodes 42 detectors per photosensor and the current Philips® PET camera decodes an average of 67 detectors per photosensor. Therefore, the current PQS detector design has 5.4-18 times more detector pixels per photosensor than current clinical PET, which translates to 12.5-76 times smaller voxels size in the tomographic image of a PQS camera, which is significant in detecting much smaller cancer lesion volume or much earlier cancer, thereby enabling better prognosis and cancer management.
The PQS detector design is typically implemented in the form of rectangular detector panels as shown in FIG. 6. However, in PET cameras, the PET detectors generally form a detection ring circumscribing the subject as shown in FIG. 7.
For SPECT cameras, although a detector ring is not needed, a large curve detector panel would allow the detector to be placed closer to the body contours (such as cardiac studies) to achieve higher SPECT resolution and sensitivity. To form a ring or curve detection system, multiple PQS detector panels can be placed next to each other to form a polygonal ring or curve, but PQS generally requires a detector-free zone (no scintillation crystal) that is the size of a half of a photosensor at each of the 4-edges of a rectangular panel as shown in FIG. 6. Hence, placing PQS panels adjacent to each other to form a detector ring or curve would cause a detector gap of one photosensor size between two detector panels. The detection gaps are not desirable for several reasons: (a) lower detection sensitivity, (b) inadequate data sampling unless the detection system rotates, which increases mechanical complexity, reliability, positioning errors and production cost. For the new “time-of-flight” PET applications, the detection gap without rotating the detection system is especially problematic in the image reconstruction process.
As shown in FIGS. 6 and 7, the scintillation blocks form detector arrays that are coupled to the PMT to form large detector panels. The panels can be placed proximal to each other to form a ring as shown in FIG. 7. As shown in FIGS. 6 and 7, the gaps between two panels on the ring are equal to or larger than one PMT size (half a PMT on each adjacent panel) because the blocks do not entirely cover the circular PMTs at the ends of the array.
For a large camera system with 12 PQS detector panels similar to that shown in FIG. 7, hundreds of photosensors detection channels (photosensors and supporting electronics) can be saved to further lower the production cost of a camera using the PQS detector design.
A continuous PET detector ring or partial ring using the PQS design can be implemented that eliminates the gaps between the panels shown in FIG. 7. As illustrated in FIG. 8, for this configuration four surfaces of each detector block must be precisely tapered to form a symmetric pentagon. The pentagonal blocks can then be butted against each other to form a ring or partial ring as shown in FIG. 7. This configuration is compatible the PQS design whereby each quadrant of each photosensor is placed adjacent to a quadrant of a scintillation crystal array.
While the configuration shown in FIG. 7 addresses some of the shortcomings of prior systems, it also results in high manufacturing costs for large scale systems.
One straightforward way to implement a PQS technology is to pack many crystal arrays/blocks into a large panelized detector module and circumscribe the imaged object by a relatively small number (e.g., four or six) such detector panels. However, there is degradation in image resolution even for the region near the center of the field-of-view and sensitivity loss caused by the gaps between two detector panels. The earlier generations of the PennPET (University of Pennsylvania) with an hexagonal detector ring demonstrated such effects, which led to the “circularization” of the later PennPET (Philips Nal PET) using curved NaI(Tl) crystals.
In one embodiment, the gap between two panels of a HOTPET camera in brain mode is 14.8 mm, and 12 such gaps cause at least a ten percent loss of the detection sensitivity. While an accurate rotation gantry can be used to rotate the apparatus and fill the gaps of missing LOR in the sinograms, such an implementation also increases the production cost.
For generally smaller detector ring designs, each detector array/block can be placed “archually” on the detector circle. To “circularize” the PQS design, which requires two adjacent arrays to share the same PMT, each detector array/block has to be ground (by a small amount) to a slightly pentagonal shape as shown in FIG. 8. On the circumferential dimension, the last two rows of crystals in each block/array also needed to be ground down to a slight taper.
In exemplary embodiments of the design shown in FIG. 8, each detector block can be manufactured from a rectangular cross-section block first by placing the block into a form (or jig) to be ground down on four surfaces to the desired shape. With tapered end-crystals in each block, all the adjacent blocks are glued together to form a solid cylindrical ring, thus providing almost one hundred percent packing fraction with the highest possible detection sensitivity.
Existing systems, including Philips® Medical System, include a close-to-gapless PET detector ring using a modular design with each detector module consisting of an array of optically isolated scintillation crystals. In certain embodiments, this crystal array is coupled to a thick (e.g., 25 mm) solid piece of light transparent material to disperse the light. This solid transparent plastic piece is curved or slanting on the output end. In such embodiments, the solid curved light guide can then be coupled to a large array of PMT. In particular embodiments using this configuration, the scintillation light is distributed to all fifteen PMT coupled to the whole detector module (though the position is determined by the nearest seven PMT), which is more than the four PMT in the PQS design. In certain embodiments, the design is essentially a conventional gamma camera Anger-logic positioning design modified from a flat camera head to a curve camera head.
In such embodiments, the light guide (a) absorbs light signal and (b) reduces the light going to the PMT by a wide area light dispersion (larger than the seven PMT used for extracting the scintillating position), thereby reducing the positioning signal strength and positioning accuracy. Furthermore, in traditional gamma camera, it is known that the probability of gamma-event “pileups” (e.g., one, two, or three events hitting the detector head during the processing of a prior event) is higher in such a large detector head with all the PMT receiving light from all the crystal pixels.